In the United States, sixty million people suffer from chronic sinusitis and allergic rhinitis and are treated by means of topically applied antihistamines, antibiotics, decongestants, and pain relievers. Many of these drugs would work more effectively in relieving symptoms if they could be applied directly to all of the affected areas. However, the devices utilized thus far to deliver these drugs have proven to be extremely inadequate, if not useless, in reaching all areas needed especially the deep nasal cavity, olfactory region, and paranasal sinuses critical to the treatment of some of these diseases and conditions. In addition to topically applied drugs (e.g., such as particular drugs in the categories listed above), there are a wide variety of systemically-absorbed drugs that are delivered intranasally. Moreover, a completely new field of nose to brain drug delivery is emerging. Current devices utilized for such systemically-absorbed drugs have also proven to be inadequate for many applications.
Current delivery systems comprise, for example, metered dose spray bottles and pneumatic (e.g., compressed air) atomizers that eject the medicine into the nostrils in large particles, or streams of atomized liquid. While a substantial mass of aerosolized particles can be quickly ejected or projected from such devices, the ejected or projected particles are relatively large, such that the efficacy of medicine administered in this manner is limited because of variable user skill and inadequate delivery and/or target distribution. For example, because of the relatively large particle sizes and the velocity vectors and characteristics of the particles, medicines delivered in this manner reach very little of the nasal mucosa and essentially no part of paranasal sinuses. Instead, such devices spray the particles into, for example, the anterior nasal cavity where the substantial mass of the particles impact the surfaces and drip out the nostril, or quickly clear along the floor of the nasal cavity. In cases of severe congestion or nasal polyps, the medicine often does not proceed beyond the nostril and has no chance of being effectively absorbed into the bloodstream in the necessary area of the nasal cavity. Therefore, while current prior art metered dose spray bottles and pneumatic atomization systems allow for rapid mass delivery, they are typically of relatively crude simplistic design, and substantially waste medicament because they do not provide adequate particle size distributions or delivery targeting for many purposes (e.g., they do not allow for particles to penetrate or reach high into the nasal cavity, and be retained therein, as required for systemic nose to brain delivery, or for paranasal sinus delivery.
As an improvement, pneumatic (e.g., compressed air) nebulizers have been developed and are familiar in the art. Fundamentally, nebulizers are distinguished from simple atomizers by the presence in the former of an ‘impaction or stagnation baffle’ placed, adjacent the compressed gas orifice, in the aerosol stream. Typically, for pneumatic nebulizers, compressed gas is delivered through a compressed air channel and orifice (jet) of a compound integrated aerosolization nozzle causing a region of negative pressure (Venturi effect) in close proximity to a restricted liquid/solution channel or capillary. The liquid to be aerosolized is entrained, by virtue of its proximity to the restricted liquid channel within the nozzle configuration, into the jet orifice gas stream and is sheared into a liquid film or ligaments that may collapse into initial droplets under the influence of surface tension. While a small proportion of the initial droplets are smaller (e.g., 5 μm or less), the predominant portion of such initial droplets and/or film/ligaments are substantially larger and are subsequently violently shattered upon impaction with the closely spaced impaction/stagnation baffle, which serves to provide for production of smaller droplets and for return of larger droplets to the liquid reservoir. For efficacy in optimizing smaller particle production, the impaction/stagnation baffle is placed extremely close to the compressed air orifice, typically within a fraction of a millimeter from the jet or nozzle orifice. Because of the close spacing, the impaction/stagnation baffle also serves to redirect compressed gas flow laterally toward the walls of the atomization chamber, and smaller particles (e.g., 5 μm or less, corresponding to both shattered and initially atomized small unshattered particles) are thereby carried laterally toward the walls of the nebulization chamber. While most of such laterally directed particles are thereafter collisionally ‘consumed’ by walls/surfaces of the atomization chamber, a small proportion of such laterally-directed particles are again redirected toward the user by the user's inhalation stream and are thereby rendered deliverable to the user (e.g., deliverable as a mist or vapor of very tiny particles to the lungs by means of a user breathing the medicine-containing particles from a pipe attachment or, in the case of young children, a face mask, e.g., inhalation of nebulized particles during an asthma attack).
Therefore, prior art closely spaced impaction/stagnation baffles provide two functions: (i) shattering of larger particles into smaller particles; and (ii) laterally redirecting smaller particles. However, in either instance, the deliverable particles do not have, upon generation, a velocity vector path in the direction of the user that is not obstructed by the impaction/stagnation baffle, and the particle velocity vectors are such that the particles thus either impact on the baffle, laterally impact on the atomization chamber wall/surfaces, or are laterally directed and subsequently directed toward the user. Significantly, therefore, with prior art nebulizers, there are no particles that have, as initially generated, velocity vectors with paths toward the user that are not obstructed by the impaction/stagnation baffle, and delivery of such particles is thus entirely dependent upon redirecting particles around the baffle by inhalation facilitated flow redirection. Significantly therefore, not only is the size range of deliverable particles limited by such designs (e.g., to those small enough to be laterally directed and redirected toward the user (e.g., 5 μm or less) by the inhalation stream, but the delivery efficiency is limited because of the small percentage of particles that avoid being ‘consumed’ on the baffle, and on the walls and surfaces of the atomization chamber because of the indirect paths that the deliverable particles must take. This is a significant limitation of prior art devices.
Fundamentally, with prior art pneumatic nebulizers, while the impaction/stagnation baffle serves to redirect the compressed air flow direction (typically at right angles to the longitudinal jet axis) and return larger droplets to the liquid reservoir for re-entrainment, the creation and size of the generated deliverable particles are entirely determined by violent impaction with the baffle subsequent to entrainment of the solution by the compressed air jet of the nozzle, and those shattered particles that don't then impact the side-walls are drawn to the user during user inhalation. Droplet size is typically reported as mean mass aerodynamic diameter (MMAD), which is the diameter around which the mass of the aerosol is equally divided; that is, the calculated aerodynamic diameter that divides the particles of an aerosol (a gaseous suspension of fine liquid or solid particles) in half, based on the mass of the particles (by mass, 50% of the particles will be larger than the MMAD and 50% of the particles will be smaller than the MMAD). Therefore MMAD is used to characterize a population of droplets produced, and does not refer to the size of individual droplets. The particle size distribution of any aerosol may thus be statistically described by the median aerodynamic diameter along with the geometric standard deviation (GSD) based on the weight and size of the particles. Significantly, it should be appreciated, that because the volume (and hence the mass) of the droplet is determined by the cube of the radius (v=4/3πr3), most of the particles will be smaller than the MMAD. The respirable dose is sometimes reported as the respirable mass, which is the output of droplets from the nebulizer in, for example, a respirable range of 1-5 um. Therefore, with prior art pneumatic nebulizers, the size and output of droplets comprising the respirable mass is entirely determined by the impaction and shattering function of the closely opposed impaction/stagnation baffle, and where a small but deliverable proportion of the laterally-directed particles avoid impacting the side-walls of the atomization chamber and are rather carried to the user in the user's inhalation stream.
Typically, a device selected for administration of pharmacologically active aerosol to the lung parenchyma should produce particle sizes with a mass median aerodynamic diameter (MMAD) of 1-3 microns. For airway deposition MMAD should be around 2-5 microns. Relatively small particle size is important for lung delivery in that, for example, it allows passage of the drug through heavily congested airways over a sufficient period (e.g., of about 10 minutes), to allow for deep lung penetration. Such nebulizers are used, for example, by asthmatics in response to an asthma attack.
With reference to FIG. 1, such prior art pneumatic nebulizers generally have, in addition to a closely opposed impaction element/baffle, a compound integrated aerosol nozzle comprising a compressed air or fluid channel with an end orifice, along with an integrated solution channel in communication with a liquid or solution (e.g., medicine solution). Moreover, such nebulizers generally correspond to one of two types; namely an ‘internal mixing’ (FIG. 1A) design or an ‘external mixing’ (FIG. 1B) design (see, e.g., Hess, D. R., Respriatory Care, 435:609-622, 2000 for a discussion of nebulizer designs incorporated herein by reference). Generally speaking, with internal mixing designs, gas flow interacts with the solution prior to leaving the nozzle exit orifice. For example, in FIG. 1A, the nozzle is concentrically mounted around a compressed gas delivery tube/channel (with end orifice) such that between the tube and nozzle there is a narrow interspace channel in communication with a liquid/solution reservoir. The exit of compressed gas from the gas delivery tube orifice causes solution to be drawn up through the restricted interspace to form an ascending stream of air and solution which leaves from the nozzle orifice and strikes the baffle to cause atomization of the particles (see also FIG. 1 of U.S. Pat. No. 6,796,513). By contrast, with external mixing, jet gas and the solution interact after both leave the nozzle. For example, in FIG. 1B, the nozzle orifice is a compound orifice, comprising a gas delivery tube/channel (with end orifice) that is coplanar with respect to a concentric solution channel orifice. In such designs fluid must leave the solution channel orifice (and the nozzle) before it can interact with the jet gas. The exit of compressed gas from the gas delivery tube orifice (and thus from the nozzle) causes solution to be drawn from the narrow solution channel and orifice (and thus from the nozzle) where it subsequently interacts with the jet gas to form a stream of air and solution which strikes the baffle to cause atomization of the particles. Different jet nebulizers have different output characteristics determined by the design of the air jet and capillary tube orifices, their geometric relationship with each other and with the closely opposed impaction baffles. In such prior art configurations, the major output determinant is generally the level/strength of the driving gas flow. So-called open ‘vented’ versions of these nebulizer designs allow for intake of ambient air during user inhalation to increase particle flow to the user and thus increase, at least to some extent, the effective nebulizer output at least during the inhalation phase.
Unfortunately, conventional jet nebulizers, including open vented versions, are highly inefficient because much of the aerosol is wasted during exhalation or excessively recycled within the nebulizer. In particular nebulizer designs, some aerosol waste is prevented by having one-way valves near the mouthpiece that redirect exhalation so that is does not substantially exhaust through the open inhalation vent in the primary aerosol generation chamber. However, even in these designs, between 93 and 99% of the primary droplets are caught on the internal baffles and structures and typically returned to the solution reservoir for re-entrainment, resulting in low output and/or protracted nebulization times. Additionally, in view of the pervasive use of restricted or narrow liquid feed channels to the medicament reservoirs means, while prior art atomizers and nebulizers are adequate for generating particles from low viscosity solutions (e.g., up to 5 centipoise), they are incapable of delivery of more viscous solutions (e.g., 5-105 centipoise). Thus, most such currently used nebulizers are not sufficiently effective at delivering enough medicament formulation (especially viscous drug solutions) in a practical or reasonable time-period because of restrictive liquid feed channels and the requirement for impaction/stagnation baffle configurations to shatter and size the particles. Additionally, even if more powerful compressor means were to be employed in such designs, there would be attendant increases in device size, weight and expense, and also (at least in particular designs) an increase the aerosol waste during exhalation phases. Moreover, increased compressed air flow would not eliminate the excessive 93 to 99% recycling of impacted medicine droplets returned to the solution reservoir from the impaction baffles. Additionally, even if there was an amount and/or quality of output sufficient for particle delivery to the lungs, absent an appropriate particle generation and dispersion means (as taught herein below by applicants), such prior art nebulizers are not effective for nasal delivery of drugs (e.g., antibiotics, etc.), because the generated particles are (i) not appropriately sized or dispersed to effectively penetrate into the nasal cavity and/or paranasal sinuses, and (ii) not delivered in a direct flow path to enable efficient delivery of sufficient quantities of medicament in a practical time-frame.
There is, therefore, a pronounced need in the art for delivery methods and devices that enable more efficient output and delivery of aerosolized particles. There is a pronounced need for devices that reduce or eliminate the dependence on baffle impaction and flow redirectioning for generation and determination of particle size, not only to reduce the extent/amount of recycling and re-entrainment of baffle-impacted solution droplets to allow for shorter, more user-friendly delivery periods, but also to provide for generation of a broader range of particle sizes to enhance dynamic output.
There is a pronounced need in the art for more effective methods and devices for delivery of aerosolized medicaments of higher viscosity.
There is a pronounced need in the art for more effective methods and devices for delivery of medicament to treat patients for certain conditions without taking the medicament orally or through the lungs. There is a pronounced need for more effective and efficient delivery to all areas of the nasal cavity and paranasal sinuses, and for more strategic or targeted delivery of medicament to specific regions of the nasal cavity, nasal olfactory region and paranasal sinuses. There is a pronounced need in the art for more effective methods and devices to effectively administer therapeutic agents systemically via the nasal passages, through the various channels from the olfactory region to the brain and the deep paranasal sinuses. There is a pronounced need for more effective methods and devices to for delivery of drugs to the brain to treat conditions of the central nervous system (CNS); that is, for ‘Nose-to-Brain’ delivery (e.g., to bypass the so-called blood brain barrier). There is a pronounced need for ocular and oral delivery using more efficient devices, and more efficient means for aerosolization and delivery of perfume, fragrance, essential oil or cosmeceutical agents and the like to the vicinity or surfaces or users or targets.